In percutaneous transluminal coronary angioplasty (PTCA), a balloon catheter is inserted through a brachial or femoral artery, positioned across a coronary artery occlusion, and inflated to compress against atherosclerotic plaque to open, by remodeling, the lumen of the coronary artery. Problems with PTCA include formation of intimal flaps or torn artery linings, both of which can create another occlusion in the lumen of the coronary artery. Moreover, thrombosis and restenosis may occur several months after the procedure and create a need for additional angioplasty or a surgical bypass operation. Stents are used to address these issues. Stents are small, intricate, implantable medical devices and are generally left implanted within the patient to reduce occlusions, inhibit thrombosis and restenosis, and maintain patency within the vascular lumens, such as, for example, the lumen of a coronary artery.
Radially expandable endoprostheses are artificial devices implanted in an anatomical lumen. An “anatomical lumen” refers to a cavity, duct, or a tubular organ such as a blood vessel, urinary tract, and bile duct. Stents are examples of endoprostheses that are generally cylindrical in shape and function to hold open and sometimes expand a segment of an anatomical lumen. Stents are often used in the treatment of atherosclerotic stenosis in blood vessels. “Stenosis” refers to a narrowing or constriction of the diameter of a bodily passage or orifice. In such treatments stents reinforce the walls of the blood vessel and prevent restenosis following angioplasty in the vascular system. “Restenosis” refers to the reoccurrence of stenosis in a blood vessel or heart valve after it has been treated by, for example, balloon angioplasty, stenting, or valvuloplasty with apparent success.
The treatment of a diseased site or lesion with a stent involves both delivery and deployment of the stent. “Delivery” refers to introducing and transporting the stent through an anatomical lumen to a desired treatment site, such as a lesion. “Deployment” corresponds to the expansion of the stent within the lumen at the treatment region. Delivery and deployment of a stent are accomplished by positioning the stent about one end of a catheter, inserting the end of the catheter through the skin into an anatomical lumen, advancing the catheter in the anatomical lumen to a desired treatment location, expanding the stent at the treatment location, and removing the catheter from the lumen. In the case of a balloon expandable stent, the stent is mounted about a balloon disposed on the catheter.
Mounting the stent typically involves compressing or crimping the stent onto the balloon prior to insertion in an anatomical lumen. At the treatment site, within the lumen, inflating the balloon expands the stent. The balloon may then be deflated and the catheter withdrawn from the stent and the lumen, leaving the stent at the treatment site. In the case of a self-expanding stent, the stent may be secured to the catheter via a retractable sheath. When the stent is at the treatment site the sheath may be withdrawn which allows the stent to self-expand.
The structure of a stent is typically composed of scaffolding that includes a pattern or network of interconnecting structural elements often referred to in the art as struts or bar arms. The scaffolding is designed so that the stent can be radially compressed (to allow crimping) and radially expanded (to allow deployment). A conventional stent is allowed to expand and contract through movement of the individual structural elements of a pattern with respect to each other.
For example, in FIG. A shows a stent 10 having an overall body shape that is hollow and tubular. The stent 10 can be formed from wires, fibers, coiled sheet, with or without openings, or a scaffolding network of rings. The stent 10 can have any geometrical configuration, such as sinusoidal or serpentine strut configuration, and are not limited to what is illustrated in FIG. A. The variation in strut patterns is virtually unlimited.
In FIG. A, the stent 10 includes many interconnecting struts 12, 14 separated from each other by openings 16. The struts 12, 14 can be made of any suitable material, such as biocompatible metal or polymer. The stent 10 has an overall longitudinal length 18 measured from opposite ends, referred to as the distal and proximal ends, 20, 22. The stent 10 has an overall body having a tube shape with a central passageway 24 passing through the entire longitudinal length of the stent 10. At least some of the struts 12 are arranged in series to form sinusoidal or serpentine ring structures 26 that encircle the central axis 32.
The terms “axial” and “longitudinal” are used interchangeably and relate to a direction, line, or orientation that is parallel or substantially parallel to the central axis 32 of the stent 10 or central axis of a cylindrical structure. The terms “circumferential” and “circumferentially” relate to a direction along a circumference of a stent or circular structure. The terms “radial” and “radially” relate to a direction, line, or orientation that is perpendicular or substantially perpendicular to the central axis 32 of a stent 10 or a central axis of a cylindrical structure.
In many treatment applications, the presence of a stent in a body may be necessary for a limited period of time until its intended function of, for example, maintaining vascular patency and/or drug delivery is accomplished. Thus, stents are often fabricated from biodegradable, bioabsorbable, and/or bioresorbable materials such that they completely erode only after the clinical need for them has ended. The term “biodegradable” relates to a compound that can be broken down by natural processes into basic components. The term “bioabsorbable” refers to the breakdown of a compound into a simpler substance or substances that are dispersed and stored within the body but not eliminated by the body. The term “bioresorbable” refers to the breakdown of a compound into a simpler substance or substances that are eliminated by the body. For example, in the case of a bioresorbable polymer the macromolecule is cleaved into low molecular mass by-products, which are eliminated from the body by biological pathways utilizing the kidneys and lungs.
The stent must be able to satisfy a number of basic functional requirements. The stent must be capable of withstanding the structural loads, for example, radial compressive forces, imposed on the stent as it supports the walls of the anatomical lumen after deployment. Therefore, a stent must possess adequate radial strength. Radial strength, which is the ability of a stent to resist radial compressive forces, is due to strength and rigidity around the circumferential direction of the stent. After deployment, the stent must adequately maintain its size and shape throughout its service life despite various forces that may come to bear on it, including the cyclic loading induced by the beating heart. In particular, the stent must adequately maintain an anatomical lumen at a prescribed diameter for the desired treatment time despite these forces. The term “treatment time” refers to the time required for the vessel to remodel, after which the stent is no longer necessary for the vessel to maintain a desired diameter.
Stents implanted in coronary arteries are primarily subjected to radial loads, typically cyclic in nature, which are due to the periodic contraction and expansion of vessels as blood is pumped to and from a beating heart. Stents implanted in peripheral blood vessels, or blood vessels outside the coronary arteries such as iliac, femoral, popliteal, renal and subclavian arteries must be capable of sustaining both radial forces and crushing or pinching loads. These stent types are implanted in vessels that are closer to the surface of the body and are particularly vulnerable to crushing or pinching loads, which can partially or completely collapse the stent and thereby block fluid flow in the vessel. A stent having suitable radial stiffness may not have sufficient crush resistance. Other design considerations for peripheral stents are the degree of bending and axial compression the stent can withstand without mechanical loss of strength or stiffness.
In addition to high radial strength, the stent must also possess sufficient flexibility to allow for crimping on a delivery device, flexure during delivery through an anatomical lumen, and expansion at the treatment site. Longitudinal flexibility is important to allow the stent to be maneuvered through a tortuous vascular path and to enable it to conform to a deployment site that may not be linear or may be subject to flexure. A stent should have sufficient toughness so that it is resistant to crack formation, particularly, in high strain regions during crimping, delivery, and deployment.
Stents are often modified to provide drug delivery capabilities to further address thrombosis and restenosis. Stents may be coated with a polymeric carrier impregnated with a drug or therapeutic substance. A conventional method of coating includes applying a composition including a solvent, a polymer dissolved in the solvent, and a therapeutic substance dispersed in the blend to the stent by immersing the stent in the composition or by spraying the composition onto the stent. The solvent is allowed to evaporate, leaving on the stent strut surface a coating of the polymer and the therapeutic substance impregnated in the polymer.
A stent and delivery system typically undergo sterilization to reduce their bioburden to an acceptable sterility assurance level (SAL). There are numerous methods of sterilizing medical devices, the most common being ethylene oxide treatment and treatment with ionization radiation such as electron beam and gamma radiation. Generally, it is desirable for the sterilization procedure to have little or no adverse affects on the performance of a sterilized article.
A polymeric stent or delivery system generally includes means for locating the stent during the percutaneous procedure. Typically, fluoroscopic visualization is used to locate markers either imbedded in the stent or delivery device.
In the prior art there are two primary types of stents: self-expanding and plastically deformed. Self-expanding stents are commonly comprised of super-elastic materials such as Nitinol. This material is known for its ability to return to its original configuration after severe deformation such as a crushing load or longitudinal bending. However, this variety of self-expanding stent has the problem of exerting a “chronic outward force (COF)” on the blood vessel supported by the stent. It is believed that a COF exerted on a blood vessel is a main contributor to high degrees of restenosis of lesions treated by a self-expanding stent even when anti-proliferative drugs are employed.
Stents that are plastically deformed to support a vessel do not suffer from COF. Stents that are plastically deformed by a balloon have the desirable quality of being deployable to the desired diameter for supporting the vessel without exerting outward forces on the vessel exceeding those required to main position of the stent in the vessel during usage. However, plastically deformed stents have the drawback of higher potential for collapsing or becoming kinked in peripheral blood vessels and are generally not used for this reason.
In the prior art, there are generally two classes of materials used to fabricate stents: metal and polymer. A stent comprised of metallic material is generally durable. A durable metallic stent has the drawbacks that it constrains the blood vessel after remodeling has occurred, is subject to frequent strut fractures, requires dual antiplatelet therapy (DAPT) to prevent subsequent thrombosis, sometimes requires revascularization, and sometimes results in the unsatisfactory outcome of late stent thrombosis which often results in patient death. However, a stent comprised of metal offers the benefits of high strength which enables a stent to be fabricated having relatively thin wall thickness and the ability to retain luminal capacity after angioplasty.
A stent comprised of polymer material is generally biodegradable, bioabsorbable, bioresorbable, biosoluble, or bioerodable. The polymeric stent is intended to remain in the body for only a limited period of time because the stent degrades, absorbs, resorbs, or erodes from the implant site. The prior art stents fabricated of bioresorbable polymers generally are comprised of struts 12,14 having a wall thicknesses comprised of a solid polymeric cross section. It is believed that biodegradable stents allow for improved healing of the anatomical lumen as compared to metal stent, which may lead to a reduced incidence of late stage thrombosis.
However, a potential shortcoming of polymer stents compared to metal stents of the same dimension, is that polymer stents typically have less radial strength and rigidity. Relatively low radial strength potentially contributes to relatively high recoil of polymer stents after implantation into an anatomical lumen. “Recoil” refers to the undesirable retraction of a stent radially inward from its deployed diameter due to radially compressive forces that bear upon it after deployment. Moreover, polymer stents of the prior art generally have twice the stent-to-artery coverage of metal stents. The higher stent-to-artery coverage may increase the probability of blocking anatomical lumen side branches or invoking an unfavorable response resulting in restenosis or thrombosis.
Another potential problem with polymer is that struts can crack or fracture during crimping, delivery, and deployment, especially with brittle polymers. Some polymers like poly (L-lactide) (“PLLA”) are stiff and strong but can exhibit a brittle fracture mechanism at physiological conditions in which there is little or no plastic deformation prior to failure. A stent fabricated from such polymers can have insufficient toughness for the range of use of a stent. In particular, during deployment of a stent comprised of a material having insufficient ductility, strut fracture can occur when the stent is expanded during deployment with a balloon catheter beyond recommended expansion limits. As a result, cracks can occur particularly in high strain regions, which can result in mechanical failure of the stent. The terms “physiological conditions” refer to conditions within the human body including, but not limited to, body temperature (approximately 37 degrees Celsius), pH, water concentration, oxygen concentration, operating pressure, etc.
A bioresorbable stent of the prior art typically has a wall thickness formed by injection molding, extrusion, or casting. These prior art stents employ polymers with relatively low molecular orientation because the molecular orientation is limited to the random molecular orientation that is imparted on the polymer when it flows through the screw inside the barrel that conveys the material to the mold or die. Instead of building strength in the polymer, these prior art process of molding or extruding can actually reduce the strength of the polymer because the polymer is subjected to relatively high shear stresses at temperatures above the melting point of the polymer which substantially reduces the molecular weight of the polymer. The resultant stents are typically much lower strength than metallic stents and, therefore, have much thicker wall thickness or wider struts to compensate for the polymeric material that is not as strong as a metallic material. Solid wall thickness polymeric stents are sometimes strengthened by radial or longitudinal plastic deformation of the precursor tubes used in manufacturing struts as described in U.S. Pat. No. 8,012,402 B2 filed 15 Apr. 2009, “Tube expansion process for semi-crystalline polymers to maximize fracture toughness.” The larger wall thickness is problematic in small diameter stents because the polymeric stents of the prior art reduce luminal capacity because the stent struts consume a significant portion of the orifice in which fluid flows. The terms “plastic deformation” refers to a process when an object is changed permanently in size or shape permanently due to an applied force.
As previously mentioned, polymeric stents are generally produced by cutting a strut pattern in a tube using a laser. A strut pattern creates the interconnecting struts separated by open spaces that define the tubular shape of the stent. The process of cutting a strut pattern in a tube wall thickness generally requires that the tube wall thickness and ovality not vary more than 5-20 microns. Since the struts are typically 75-150 microns in width, it is necessary to use tubes having very little dimensional variation or it is impossible to accurately cut a strut pattern in a tube achieving these narrow strut widths. The prior art methods of have difficulty producing tubes meeting these stringent tolerances, especially when manufacturing a tubular precursor construct having a length greater than 15 cm.
Stents produced from polymer tubes having a wall thickness comprised of filaments, fibers, fibers, and the like can suffer from insufficient stiffness. Mechanical tests suggest that prior art polymeric stents experience less recovery of diameter when expanded by balloon after crimping. The polymeric stents are also more difficult to plastically deform and to properly size in lumen after deployment. Researchers also found that polymeric stents of the prior art have less net gain in luminal capacity at periodic checkpoints after intervention than durable metallic stents. Many polymers used for stents are relatively brittle under physiological conditions. Polymers are susceptible to mechanical instability such as fracturing while in the body. Bioresorbable stents sometimes fracture when expanded more than 17% during deployment. Bioresorbable stents are difficult to see during delivery and deployment because the polymers are not radiopaque, which leads to malposition of the stent.